Regenerative medicine aims to restore, maintain or improve function of damaged organs. Regenerative medicine bridges the gap between artificial devices and natural organs by providing biological scaffolds, which not only sustain the function of the diseased organ but also accelerate the process of regeneration. The choice of scaffold material and architecture are key to successful regeneration. Natural as well as synthetic scaffolds are routinely employed. Natural materials such as polypeptides, polysaccharides, nucleic acids, hydroxyapatites, or their composites offer excellent physiological activities such as selective cell adhesion (e.g., collagen and fibrin), mechanical properties similar to natural tissues (e.g., animal heart valves and blood vessels), and biodegradability (e.g., gelatin and chitin). However, biological materials have several disadvantages including risk of viral infection, antigenicity, unstable material supply, and deterioration, which accompanies long-term implantation. In addition, naturally-derived materials offer limited versatility in designing an exogenous extracellular matrix with specific properties (e.g., porosity and mechanical strength). Synthetic materials, by contrast, can be manufactured reproducibly on a large scale, and can also be processed into an exogenous extracellular matrix in which the macrostructure, mechanical properties, and degradation time can be readily controlled and manipulated. Exogenous extracellular matrices fabricated by biodegradable polymers eventually break down in the body, avoiding chronic foreign-body responses (Rosso, et al. (2005) J. Cell. Physiol. 203:465-70).
An ideal scaffold provides a three dimensional shape for cells to attach and a porous architecture (Bryant & Anseth (2001) J. Biomed. Mat. Res. A 59:63-72). In addition, the scaffold should be biodegradable, biocompatible and not elicit an immune response or rejection when implanted in the body. Poly (lactic acid) (PLA), poly(glycolic acid) (PGA), and their copolymers poly(lactic acid-co-glycolic acid) (PLGA) are routinely used as polymers for scaffold designing because of their wide biodegradability and biocompatibility (Chen & Ma (2006) Biomaterials 27:3708-15). Moreover, co-polymerization of polyethylene glycol (PEG) with PLA and PLGA has been used to increase the biodegradability of PEG (Bryant, et al. (2004) Biotechnol. Bioeng. 86:747-55).
Scaffolds can be classified into two categories based on their porosity, non-porous scaffolds (Fukuda, et al. (2006) Biomaterials 27:5259-67) and porous scaffolds (Norman & Desai Ann. Biomed. Eng. 34:89-101). Commonly used methods for creating unordered porous structures include polymer demixing (Dalby, et al. (2002) Tissue Eng. 8:1099-108), phase separation (Smith & Ma (2004) supra), foaming (Chen, et al. (1999) J. Biomed. Mater. Res. 44:53-62), colloidal lithography (Dalby, et al. (2004) Biomaterials 25:5415-22; Denis, et al. (2004) Langmuir 20:9335-9), self-assembly (Tu & Tirrell (2004) Adv. Drug Deliv. Rev. 56:1537-63), chemical etching (Tu & Tirrell (2004) supra; Thapa, et al. (2003) Biomaterials 24:2915-26), and thermogellation (US Patent Application No. 20060257378). Ordered structures can also be obtained by photolithography, electron beam lithography (Curtis, et al. (2001) Biophys. Chem. 94:275-83; Smith & Ma (2004) Colloids Surf. B Biointer. 39:125-31), and electrospinning (Williams, et al. (2005) Biomaterials 26:1211-8; Jayaraman, et al. (2004) J. Nanosci. Nanotechnol. 4:52-65). In addition, US Patent Application No. 20080193536 discloses microfabrication of cell-laden hydrogels using soft lithographic techniques. However, macroscopic, non-porous hydrogels, prepared for example by photopolymerization techniques, have the disadvantages of photoinitiator toxicity (Williams, et al. (2005) Biomaterials 26:1211), limited diffusion, and limited cell-cell interactions within the scaffolds (Nuttelman, et al. (2004) J. Biomed. Mater. Res. A 68:773).
In comparison, porous scaffolds with a pore size of greater than 100 μm are reported to promote bone ingrowth within the pores of the scaffolds (Langer & Vacanti (1993) Science 260:920) and promote cell-cell interactions. Insufficient interconnectivity in biodegradable hydrophobic polymers, e.g., poly(lactide-co-glycolide) copolymers, has been shown to be the limiting factor in cell colonization and new tissue formation (Crane, et al. (1995) Nat. Med. 1:1322; Ishaug, et al. (1997) J. Biomed. Mater. Res. 36:17). Microchannel conduits have been engineered in non-porous PEGDA hydrogels for generating vascularized adipose tissue grafts (Stosich, et al. (2007) Tissue Eng. 13:2881). This study highlighted the necessity of a porous scaffold structure for vascularization. The presence of interconnected pores within the scaffold was shown to enable vascular capillary ingrowth from the host tissue into the interconnected network of the scaffold (Hutmacher (2000) Biomaterial 21:2529).
One of the limiting factors of hydrogels has been the rather slow swelling property of dried hydrogels. For the dried hydrogels to swell, water has to be absorbed into the glassy matrix of the dried hydrogels. The swelling kinetics of the dried hydrogels thus depend on the absorption of water occurring by a diffusional process and the relaxation of the polymer chains in the rubbery region. Equilibrium swelling of dried hydrogels in an ordinary tablet size (e.g., 1 cm in diameter×0.5 cm height) usually takes at least several hours, and this may be too slow for many applications where fast swelling is essential. Superporous hydrogels (SPHs) are a class of macroporous hydrogels developed for fast-swelling applications (Gemeinhart, et al. (2000) Polymers Adv. Technol. 11:617). SPHs can be prepared from addition monomers, such as PEGDA, by a gas foaming technique wherein the foaming and gelation processes are simultaneous to yield hydrogels with a macroporous network. The gas foaming technique is carried out by adding macromer, initiator, and foam stabilizer to a tube; acidifying this solution to retard the polymerization process; and subsequently adding sodium bicarbonate to generate carbon dioxide bubbles, which makes the foam rise. The addition of sodium bicarbonate increases the pH, resulting in faster polymerization of macromers. Due to this methodology, SPHs have a highly porous interconnected structure and large surface-to-volume ratio throughout the scaffold (Gemeinhart, et al. (2000) supra).